Over the past few decades, the practice of medicine has been transformed by development of new medical imaging technologies, such as magnetic resonance imaging (“MRI”). MRI can provide detailed pictures of the interior of a living body, reducing the need for exploratory surgery and providing tools for more accurate and timely diagnosis of disease or injury. Besides imaging relatively static tissue, new MRI techniques have been, and continue to be, developed that allow imaging of dynamic activity in the body, such as heart function, brain activity, or blood flow.
MRI is based on a physical phenomenon known as nuclear magnetic resonance (“NMR”) which relates to the quantum mechanics of nuclear spin. Any atomic nucleus with an odd number of neutrons, an odd number of protons, or both, has a net magnetic moment, making such an atomic nucleus “NMR active” and usable for MRI imaging. In practice, 1H (“proton”) nucleus is the most widely used for imaging of the human body. The 1H (“proton”) nucleus is attractive because all living organisms contain an abundance of hydrogen and associated 1H nuclei, and because the properties of 1H give rise to relatively strong NMR signals. FIGS. 1–4 illustrate some of the basic physics exploited by NMR and MRI devices in the prior art.
In the absence of a magnetic field {overscore (B)}0 in other words when |{overscore (B)}0|=0 as shown in FIG. 1(a), the magnetic moments {overscore (m)}i of NMR-active nuclei will be randomly oriented in all directions. Because the magnetic moments {overscore (m)}i are all randomly oriented in FIG. 1(a), the magnetic moments of all the nuclei in FIG. 1(a) cancel each other out, so that the aggregate magnetic moment is zero.
When a static magnetic field {overscore (B)}0 is applied, in other words when |{overscore (B)}0|≠0 as shown in FIG. 1(b), the magnetic moments {overscore (m)}i of the NMR-active nuclei tend to align parallel or anti-parallel to the direction of the magnetic field {overscore (B)}0. A slightly higher proportion of the NMR-active nuclei align parallel to the magnetic field, so that the magnetic moments {overscore (m)}i of all the NMR-active nuclei in FIG. 1(b) do not cancel each other out in aggregate. Instead, the aggregated magnetic moments {overscore (m)}i of all the NMR-active nuclei in FIG. 1(b) form an aggregate net magnetization {overscore (M)} as shown in FIG. 1(c).
The aligned nuclei in the static magnetic field exhibit a resonance phenomenon at a specific resonance frequency determined by the Larmor relationship [Eq. 1], where f is the resonance frequency of the nucleus in the static magnetic field, γ is a proportionality constant that depends on the nucleus, and |{overscore (B)}0| is the magnitude of the static magnetic field. For 1H, the proportionality constant γ is 42.58 MHz/T so a 1H nucleus in a 2T static magnetic field has a resonant frequency f=85.16 MHz.f=γ|{overscore (B)}0|  [Eq. 1]
Because of this resonance phenomenon, if a 1H nucleus in a 2T static magnetic field is subjected to an oscillating magnetic field {overscore (B)}excite at 85.16 MHz, that 1H nucleus will absorb energy from the oscillating magnetic field {overscore (B)}excite. Such an oscillating magnetic field {overscore (B)}excite orthogonal to the static magnetic field {overscore (B)}0 can be provided, for example, by passing an oscillating current i(t) at 85.16 MHz through an excitation coil, as shown in FIG. 2. This process is known as “excitation,” and after absorbing energy the nucleus is said to be “excited.” The excitation energy will affect a portion of the NMR-active nuclei in two ways: it will tip the magnetization vector {overscore (m)} of some of the nuclei 90° and it will flip the magnetization vector of other nuclei 180°.
The 90°-tipped nuclei have a magnetization vector {overscore (m)} with a “transverse” component mxy (in the plane perpendicular to the static magnetic field {overscore (B)}0 in the z direction) and a “longitudinal” component mz (parallel to the static magnetic field {overscore (B)}0) as shown in FIG. 2. The tipped magnetization vectors {overscore (m)}i of these 90°-tipped nuclei precess in phase around the static magnetic field {overscore (B)}0 at the resonance frequency f.
If a signal detection coil is placed nearby as shown in FIG. 2, the aggregated in-phase precession of the 90°-tipped nuclei results in a time-varying magnetic field that induces a corresponding signal voltage s(t) in the signal coil at the resonance frequency f. When the excitation from the oscillating magnetic field {overscore (B)}excite is removed, the signal voltage s(t) does not stop instantaneously, but instead it decays exponentially with a time constant T2 as the 90°-tipped magnetization vectors of the nuclei fall out of phase with one another and the nuclei return to their pre-excitation state.
In some prior art MRI and NMR devices, the same coil is used both for excitation and for measurement. That is, the same coil can be used first to carry an excitation current to create the oscillating magnetic field, then the excitation current can be turned off so that the same coil can be used to measure the induced signal voltage. However, this is not required and separate coils can be used for excitation and signal detection as shown in FIG. 2.
The 180°-tipped nuclei have a magnetization vector which has been completely flipped 180°, so that these magnetization vectors do not include any transverse components. The effect of flipping some of the nuclei 180° is to reduce the aggregate magnetization {overscore (M)} compared to what it would be without the excitation field {overscore (B)}excite. When the excitation field {overscore (B)}excite is removed, the reduced aggregate magnetization {overscore (M)} does not recover instantaneously. Instead, {overscore (M)} increases exponentially with a time constant T1 until it returns to its equilibrium state in the absence of the excitation field {overscore (B)}excite.
The environment surrounding the nuclei, for example the chemical composition of the tissues containing the nuclei, as well as the relative concentration of the NMR-active nuclei being excited, affects how the tissue reacts to the excitation. For example, the time constant T2 of the decay of the transverse components in the 90°-tipped nuclei is different for different types of tissues. Similarly, the time constant T1 of the return to normalcy of the 180°-tipped nuclei depends on the material containing the nuclei. By measuring the T1 and T2 properties of the tissue at a given position, it is possible to differentiate, for example, healthy tissue from cancerous tissue, or fat from muscle.
One-dimensional spatial localization in MRI can be accomplished by varying the magnitude of the static magnetic field |{overscore (B)}0| in that dimension across the region being imaged, as shown in FIG. 3. For example, the magnitude of the static magnetic field |{overscore (B)}0| may not be uniform, but instead it may increase linearly in a region as a function of position x according to [Eq. 2], where B(x) is the magnitude of the static magnetic field at position x, B0 is the magnitude of the static magnetic field at position x=0, and G is the magnetic field gradient across the region:B(x)=B0+Gx  [Eq. 2]
By combining [Eq. 1] and [Eq. 2], it follows that a linear variation of the magnitude of the static magnetic field in the region according to [Eq. 2] results in a corresponding variation of the resonant frequency of NMR-active nuclei in that region according to [Eq. 3]:f(x)=γ(B0+Gx)  [Eq. 3]
For example, if B0=2T, and G=0.2T/m, then the resonant frequency of the 1H nucleus in the 2T static magnetic field at position x=0 m will be 85.16 MHz. The resonant frequency of the 1H nucleus in the 2.2T static magnetic field at position x=1 m will be higher, 93.68 MHz, because of the increased static magnetic field at that position. Although the 1H nucleus is used in this example calculation, this effect is not specific to any particular NMR-active nucleus, and the resonant frequency of other NMR-active nuclei can be varied in a similar fashion.
This variation in the resonant frequency of the NMR-active nuclei can enable selective excitation as a function of position. For example, if an oscillating magnetic field at a single frequency of 85.16 MHz is applied to the region, only those nuclei at position x=0 will be excited. Because only those nuclei at position x=0 will be excited, the corresponding induced voltage in a coil will have frequency components at that single frequency and will come from those nuclei at that specific position.
Similarly, if an oscillating magnetic field having components at all frequencies across the range 85.16 MHz through 93.68 MHz is applied to the region, all the nuclei from position x=0 m to position x=1 m will be excited. The resulting induced voltage in a coil will have frequency components at all frequencies across the range 85.16 MHz through 93.68 MHz since this induced voltage will come from nuclei at all positions between x=0 m and position x=1 m. By measuring the frequency content of the induced coil, for example using a spectrum analyzer or a bandpass filter, the magnitude of the induced coil voltage at a specific position or range of positions can be detected, as illustrated in FIG. 4.
Other techniques known in the art, such as phase encoding or selective excitation in a particular slice, provide localization in two and three dimensions. No matter how many dimensions are used, however, it can be seen that the spatial resolution of the system depends at least in part on the selectivity of a spectrum analyzer or narrow bandpass filter. For example, if a narrow bandpass filter or spectrum analyzer is able to select 100 different frequencies uniformly distributed in the range 85.16 MHz through 93.68 MHz, it follows that the spatial resolution of the frequency-encoding system will be approximately 1/100 m, or 1 cm.
Thus, there is a need for methods and devices able to detect alternating magnetic fields in extremely narrow frequency ranges, to thereby provide improved spatial resolution of MRI devices. What is further needed is a magnetic field sensor having improved sensitivity and a lower noise floor compared to inductive coil sensors. What is further needed is a magnetic field sensor which minimizes the amount of metal in the vicinity of the MRI magnetic field.